Computed Tomography (CT) is a known approach for obtaining images of the internal structure of a subject from its projections. CT has been shown to be an advancement compared to conventional radiography in the detection of a wide variety of diseases because of the greater contrast it achieves. However, CT involves a considerably greater amount of radiation than conventional radiography. See R. Lecomte, P. Berard, Cadorette, D. Lapointe, “Method and system for low radiation computer tomography”, U.S. Publication No. 2008/0317200 A1, Dec. 25, 2008.
Typically, in radiographic imaging systems, such as x-ray and computed tomography (CT), an x-ray source emits x-rays toward a subject or object, such as a patient or a piece of luggage. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-rays. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image. See K. C. Burr, J. J. Shiang, A. J. Couture, “Multi-layer Radiation detector assembly”, U.S. Publication No. 2010/0102242 A1, Apr. 29, 2010.
Conventional CT imaging systems utilize detectors that convert radiographic energy into current signals that are integrated over a time period, then measured and ultimately digitized. A drawback of such detectors however is their inability to provide data or feedback as to the number and/or energy of photons detected. During image reconstruction, data as to the number and/or energy of photons detected can be used to distinguish materials which appear identical in images reconstructed from conventional systems that do not provide this additional information. That is, conventional CT detectors have a scintillator component and photodiode component wherein the scintillator component illuminates upon reception of radiographic energy and the photodiode detects illumination of the scintillator component and provides an electrical signal as a function of the intensity of illumination. A drawback of these detectors is their inability to provide energy discriminatory data or otherwise count the number and/or measure the energy of photons actually received by a given detector element or pixel. That is, the light emitted by the scintillator is a function of the number of x-rays impinged as well as the energy level of the x-rays. Under the charge integration operation mode, the photodiode is not capable of discriminating between the energy level or the photon count from the scintillation. For example, two scintillators may illuminate with equivalent intensity and, as such, provide equivalent output to their respective photodiodes. Yet, the number of x-rays received by each scintillator may be different as well as the x-rays' energy, but yield an equivalent light output.
In attempts to design scintillator based detectors capable of photon counting and energy discrimination, detectors constructed from scintillators coupled to either avalanche photodiodes (APDs) or photomultipliers have also been used. However, there are varying problems that limit the use of these detectors. In the case of APDs, there is additional gain need to enable photon counting, but with associated gain-instability noise, temperature sensitivity, and other reliability issues. In the case of photomultiplier tubes, these devices are too large, mechanically fragile, and costly for high resolution detectors covering large areas as used in CT. As such, these photomultiplier tubes have been limited to use in PET or SPECT systems.
To overcome these shortcomings, energy discriminating, direct conversion detectors capable of not only x-ray counting, but of also providing a measurement of the energy level of each x-ray detected have been used in CT systems. A drawback of direct conversion semiconductor detectors, however, is that these types of detectors cannot count at the x-ray photon flux rates typically encountered with conventional CT systems. That is, the CT system requirements of high signal-to-noise ratio, high spatial resolution, and fast scan time dictate that x-ray photon flux rates in a CT system be very high, e.g. at or greatly exceeding 1 million x-rays per sec per millimeter squared.
Also, the count rate in a single detector pixel, measured in counts per second (cps) and determined by the flux rate, the pixel area, and the detection efficiency, is very high. The very high x-ray photon flux rate causes pile-up and polarization. “Pile-up” is a phenomenon that occurs when a source flux at the detector is so high that there is a non-negligible possibility that two or more x-ray photons deposit charge packets in a single pixel close enough in time so that their signals interfere with each other. Pile-up phenomenon is of two general types, which result in somewhat different effects. In the first type, the two or more events are separated by sufficient time so that they are recognized as distinct events, but the signals overlap so that the precision of the measurement of the energy of the later arriving x-ray or x-rays is degraded. This type of pile-up results in a degradation of the energy resolution of the system.
In the second type of pile-up, the two or more events arrive close enough in time so that the system is not able to resolve them as distinct events. In such a case, these events are recognized as one single event having the sum of their energies and the events are shifted in the spectrum to higher energies. In addition, pile-up leads to a more or less pronounced depression of counts in high x-ray flux, resulting in detector quantum efficiency (DQE) loss.
Direct conversion detectors are also susceptible to a phenomenon called “polarization” where charge trapping inside the material changes the internal electric field, alters the detector count and energy response in an unpredictable way, and results in hysteresis where response is altered by previous exposure history. This pile-up and polarization ultimately leads to detector saturation, which as stated above, occurs at relatively low x-ray flux level thresholds in direct conversion sensors. Above these thresholds, the detector response is not predictable and has degraded dose utilization that leads to loss of imaging information and results in noise and artifacts in x-ray projection and CT images. In particular, photon counting, direct conversion detectors saturate due to the intrinsic charge collection time (i.e. dead time) associated with each x-ray photon event. Saturation will occur due to pulse pile-up when x-ray photon absorption rate for each pixel is on the order of the inverse of this charge collection time.
Previously conceived approaches to enable photon counting at high x-ray flux rates include using bowtie shaped filters to pre-shape the profile of the flux rate along the detector, compensating for the patient shape and producing a smaller dynamic range of flux across the field of the detector. What can be problematic, however, is that the bowtie filter may not be optimal given that a subject population is significantly less than uniform and can vary in shape. In such cases, it is possible for one or more disjointed regions of saturation to occur or conversely to over-filter the x-ray flux and create regions of very low flux. Low x-ray flux in the projection will ultimately contribute to noise in the reconstructed image of the subject.
Another proposed approach to accommodate high flux rates has been to subdivide the pixel into multiple sub-pixels, each sub-pixel connected to its own preamplifier. By reducing the area of the direct conversion sub-pixel, the flux rate capability may be increased as fewer photons are collected in the smaller area. However, the signal-to-noise ratio of the resulting signal may be reduced, and the level of cross-talk will be disadvantageously significant due to the increased perimeter between sub-pixels. Crosstalk in a direct conversion detector takes the form of charge sharing between pixels for x-rays that are absorbed near the boundaries between pixels. Charge sharing may cause the photon to be missed entirely or mislabeled in energy. In either case, the DQE is decreased and spectral response is of reduced fidelity as a result of using subdivided pixels, each connected to its own amplifier.
A photon counting output readout for solid-state photomultiplier devices is available in a few prototype circuits based on the typical approach with scintillator coupling. Several patents in this area have been issued to the group of K. Burr and J. Short at General Electric Company, for example, referring to J. D. Short, G. E. Possin, J. W. Le Blanc, R. G. Rodrigues, K. C. Burr, A. J. Couture, W. Li, “Photon counting CT Detector Using Solid-State Photomultiplier And Scintillator”, U.S. Pat. No. 7,403,589 B1.
When a solid-state photomultiplier device is used in a particular CT readout approach, the pile-up phenomenon is present at the output signal with a duration of about 50-200 ns, thus pulses generated by distinct received photons are overlapped and this may cause saturation of the respective amplifier.